Ultrasonic transducers are particularly useful for non-invasive as well as in vivo medical diagnostic imaging. Conventional ultrasonic transducers are typically fabricated from piezoelectric ceramic materials, such as lead zirconate titanate (PZT), or PZT-polymer composites, with the transducer material being diced or laser cut to form a plurality of individual elements arranged in one-dimensional or two-dimensional arrays. Acoustic lenses, matching layers, backing layers, and electrical interconnects (e.g., flex cable, metal pins/wires) are typically attached to each transducer element to form a transducer assembly or probe. The probe is then connected to control circuitry using a wire harness or cable, where the cable contains individual wires to drive and receive signals from each individual element. An important aim of ongoing research in ultrasonic transducer technology is increasing transducer performance and integrability with control circuitry while decreasing transducer size, power consumption and signal loss due to the cabling. These factors are particularly important for two-dimensional arrays required for three-dimensional ultrasound imaging.
The production of ever-smaller transducers is facilitated by micromachining techniques. There are two types of micromachined ultrasonic transducers (MUTs): capacitive MUTs (cMUTs) and piezoelectric MUTs (pMUTs). cMUTs operate by electrostatically actuating a suspended surface micromachined membrane via two opposing electrodes. Acoustic pressure is generated by vibrating the membrane, and received signals are measured as the membrane deflects proportional to the acoustic energy reflected back. pMUTs generate or transmit ultrasonic energy through application of ac voltage to the piezoelectric material causing it to alternately expand and contract, thereby flexing or vibrating the membrane. Received ultrasonic energy generates electrical charge in the piezoelectric layer due to vibrations of the bulk micromachined membrane.
Because pMUTs have a higher energy transduction mechanism, the piezoelectric layer, they generally have higher ultrasonic power capability than cMUTs. Thus, pMUTs transmit more ultrasonic energy and are more sensitive in receive for smaller element sizes compared to cMUTs. Elements in pMUT arrays also have higher capacitance (on the order of 100-1000 pF), so element impedance is lower and impedance mismatch to the cabling and electronics is less of an issue than for cMUT elements with capacitance on the order of 1 pF.
cMUTs can be energized by applying appropriate dc and ac voltage signals to the electrodes, such that an appropriate ultrasonic wave is produced. The dc voltage is required to electrostatically pull the membrane close to the substrate surface, thereby reducing the dielectric air gap, and the ac voltage vibrates the membrane to produce acoustic energy. Similarly, when electrically biased with dc voltage, the membrane of the cMUT may be used to receive ultrasonic signals by capturing reflected ultrasonic energy and transforming that energy into movement of the electrically biased membrane, which then generates a voltage signal. Another advantage of pMUTs is that they do not require the large (>100V) dc bias voltage for operation in addition to the ac signal. Lower ac voltages (<50V) are applied to activate the piezoelectric vibration for transmit, and receive signals are generated by the received ultrasonic energy alone (no applied voltage is required). One advantage of cMUTs is their higher bandwidth (>100%) over pMUTs (typically <50%) which provides higher frequency range of operation. This is beneficial for optimizing imaging resolution in different parts of the body which requires different frequency ranges.
A major advantage of MUTs is that they can be miniaturized and directly integrated with control circuitry. cMUTs with through-wafer via connections can be made by etching vias in a silicon wafer, coating the wafer with a thermal silicon dioxide for insulating regions and with polysilicon for electrical contacts, and then building up the cMUT membrane elements on the top surface of the wafer. Metal pads and solder bumps are deposited on the bottom surface of the wafer in order to solder the cMUT chip to semiconductor device circuitry. One disadvantage of such a device is that relatively high resistivity polysilicon, compared to metals, is used as the conductive material in the vias. Because of the very low signal strength (on the order of several mV or less) generated by cMUTs in the receive mode, the signal to noise ratio can be problematic during operation of the cMUT. Also, the low capacitance of cMUT elements produces high impedance, and therefore impedance mismatch with the electronics and cabling are greater which contributes to increased signal loss and noise. High resistance in the through-wafer vias further exacerbates the high element impedance problem. In addition, significant resistance in the vias will cause more power consumption and heat generation during operation when applying drive signals to cMUTs for transmit.
Another disadvantage of the cMUT device with polysilicon through-wafer interconnects is the processing temperature of forming the thermal silicon dioxide insulator and the polysilicon conductor. Processing temperatures for these steps are relatively high (600-1000° C.), thus creating thermal budget issues for the rest of the device. Because of these processing temperatures, the cMUT elements must be formed after the through-wafer vias are formed, and this sequence creates difficult processing issues when trying to perform surface micromachining on a substrate with etched holes through the wafer.
MUTs formed with through-wafer interconnects can be combined with control circuitry, thereby forming a transducer device, which can then be further assembled into a housing assembly including external cabling to form an ultrasonic probe. The integration of MUTs with control circuitry may significantly reduce the cabling required in the ultrasonic probe. The ultrasonic probe may also include various acoustic lens materials, matching layers, backing layers, and dematching layers. The housing assembly may form an ultrasonic probe for external ultrasound imaging, or a catheter probe for in vivo imaging.
Previously, joining a conventional ceramic ultrasonic transducer to electrical control circuitry required the use of many individual wires to connect each transducer element to the control circuitry. In the case of large transducer arrays, especially two-dimensional arrays having hundreds or more elements, large wiring harnesses were required. Large wiring harnesses drive up the cost and size of the ultrasonic probe, also making the probe difficult to manipulate by the user and impractical for use in catheter applications. Thus, it is desirable to reduce the cost and size of ultrasonic probes, especially for use in vivo.
One way of reducing the size of ultrasonic probes is to form the control circuitry on an integrated circuit assembly and attach the transducer directly to the integrated circuit.